The field of the invention is positron emission tomography (PET) scanners, and particularly gamma ray detectors employed to locate and count the positron annihilation events.
Positrons are positively charged electrons which are emitted by radionuclides that have usually been prepared using a cyclotron. The radionuclides most often employed in diagnostic imaging are fluorine-18 (18F) , carbon-11 (.sup.11 C) , nitrogen-13 (.sup.13 N), and oxygen-15 (.sup.15 O). These are employed as radioactive tracers called "radiopharmaceuticals" by attaching them to substances, such as glucose or carbon dioxide. The radiopharmaceuticals are injected in the patient and become involved in such processes as blood flow, fatty acid and glucose metabolism, and protein synthesis.
As the radionuclides decay, they emit positrons. The positrons travel a very short distance before they encounter an electron, and when this occurs, they are annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to PET scanners--each gamma ray has an energy of 511 keV and the two gamma rays are directed in nearly opposite directions. An image is created by determining the number of such annihilation events at each location within the field of view.
The PET scanner includes one or more rings of detectors which encircle the patient and which convert the energy of each 511 keV photon into a flash of light that is sensed by a photomultiplier tube (PMT). Coincidence detection circuits connect to the detectors and record only those photons which are detected simultaneously by two detectors located on opposite sides of the patient. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes hundreds of millions of events are recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well known computed tomography techniques.
The resolution of the reconstructed image is determined in part by the number and size of the detector elements which encircles the patient. By decreasing the size of each detector and increasing their number, the image resolution is increased, but the cost and complexity of the resulting electronic circuitry is also increased. One method which is used to improve resolution without driving the costs up is to employ a set of separate scintillation crystals of small size with each electronic circuit and encode the signal to indicate which crystal recorded the scintillation event. In U.S. Pat. No. 4,394,576, for example, four separate crystals are employed with two PMTs and the relative magnitudes of the signals detected by the PMTs indicates which of the four crystals recorded the event. A single channel thus produces a crystal address which has four times the resolution of prior detectors that employ one crystal with each PMT. This method is not effective if more than four crystals are used with each pair of PMTS, and in an effort to further improve image resolution, U.S. Pat. Nos. 4,743,764; 4,749,863 and 4,750,972 disclose a structure which employs a single large scintillation crystal with four PMTS. Slots are cut in the large crystal to divide it into separately identifiable segments, and by varying the depth of these slots, the location of the segment in the crystal which recorded the scintillation event can be identified. As many as eight segments in a row can be separately identified using this technique, but it requires the use of large, expensive crystals, and the slots have a thickness which reduces the packing fraction of the detector.